<A> Wavelength-Tunable Light Generator and Optical Coherence Tomography
(1) Optical Coherence Tomography (OCT)
Optical coherence tomography (OCT) that utilizes low-coherence light is a new medical measurement technique that permits observation of a tomogram close to the surface of living bodies at a resolution on the order of several tens of μm. OCT has already been practically used in the clinical observation of eye tissue and makes it possible to perform tomographic observation of eye tissue lesions (for example, detachment of the retina) with microscopic accuracy (See Non-Patent Document 1, for example). Clinical applications of this technique have started, but further development of tomographic observation within living bodies combined with an endoscope and so forth are expected.
OCT, which is being practically used now, utilizes a measurement technique known as ‘Optical Coherence Domain Reflectometry’ (OCDR) that requires mechanical scanning. Meanwhile, research on techniques known as Frequency Domain (FD)-OCT and Optical Frequency Domain Reflectometry (OFDR)-OCT=began recently, which do not require mechanical scanning. Each of these techniques will be described hereinbelow. Further, although referred to as OFDR-OCT in previous documents, because this technique is also called FD-OCT in recent documents, FD-OCT, which is the recent term, will be used in the description that follows. The OCT of the present invention is similar to FD-OCT in that measurements are done in optical frequency regions. Therefore, the name OFDR-OCT is subsequently used for the OCT of the present invention to differentiate it from FD-OCT.
(2) OCDR-OCT
The measurement principle of OCDR-OCT involves using a Michelson interferometer with a low-coherence light source to measure the optical path-length that is rendered as a result of measurement light 2 being irradiated into a sample (living bodies, for example) 1, reflected or backscattered at tissue boundaries 3 within the sample 1, then re-emitted from the sample 1, as shown in FIG. 7. Subsequently, in order to simplify the description of ‘reflection’ or ‘backscattering’, these are also simply referred to as ‘reflection’.
A portion of the light 2 that enters the sample 1 is reflected as a result of the difference in the refractive indices of tissues on both sides of the tissue boundaries 3 and is re-emitted from the sample 1. The structure in the depth direction within the sample 1 can be found by measuring the optical path-lengths through which re-emitted light 4 has passed. Here, the position of the surface of the sample 1, used as the depth reference point, is provided by reflected light from the surface. Therefore, a cross-sectional image or three-dimensional image or the like of the inside of the sample 1 can be obtained by scanning the entry position of the measurement light 2 on the surface of the sample 1.
FIG. 8 is a schematic view of an OCDR-OCT device. As shown in FIG. 8, light emitted from a light source 5, usually a superluminescence diode (also referred to as ‘SLD’ hereinafter) is used, is input to a Michelson interferometer 6. This light is divided by a beam splitter 7 and one of the light components is made to converge in the form of a narrow beam and irradiated into the sample 1. The other divided light component is irradiated onto a reference mirror 8. Each of the light components is reflected by the sample 1 and the reference mirror 8, respectively. They are combined by the beam splitter 7 and the combined light enters a photodetector 9.
An SLD has a broad spectral width of approximately 20 nm and, therefore, the coherence length of the emitted light is short at several tens of μm. For example, the coherence length of SLD light with a center wavelength of 850 nm and a spectral width of 20 nm is 15 μm. Hence, the signal light 10 and reference light 11 only interfere with each other when the optical path-lengths of the signal light 10 and reference light 11 agree within the short coherence length range. That is, when the reference mirror 8 is scanned in the direction of the optical axis of the reference light 11, the output of the photodetector 9 shows an interference pattern 15 (called an ‘interferogram 15’ hereinbelow), as shown in FIG. 9, with a width on the order of the coherence length, only within the short distance 14 where the optical path-length of the signal light 10 and reference light 11 match. In FIG. 9, the vertical axis 12 represents the output of the photodetector 9 and the horizontal axis 13 represents the displacement-distance of the reference mirror 8. The optical path-length of the signal light 10 can be directly found from the position of the reference mirror 8 where the interferogram 15 appears.
The resolution of this method is determined by the coherence length of the light source used and is typically on the order of 10 to 15 μm. Further, the time required for a single measurement is determined by the time required for scanning of the reference mirror 8 and is typically on the order of one second (See Non-Patent Document 1, for example) even for fast measurements.
(3) FD-OCT
The occurrence of mechanical vibrations due to the requirement for the mechanical scanning of the reference mirror 8 is unavoidable in OCDR-OCT and there are restrictions on the scanning distance at a high speed and also on the scanning speed. Because the scanning speed is restricted, there is the problem that the sample (biological sample, for example) must be constrained during measurement, and so forth. As a result, tomographic applications to tissues other than eye tissue, which is relatively easy to constrain, are not straightforward.
As a method not requiring scanning of the reference mirror 8, a frequency domain FD-OCT has been proposed (Non-Patent Document 2, for example), in which a diffraction grating 21 and a charge-coupled device (CCD) 16 are arranged on the output side of the Michelson interferometer as shown in FIG. 10. The spectrum of the output light is measured by the CCD 16, while the reference mirror 8 remains fixed, and the interferogram is calculated and constructed from the spectrum.
The principles of the FD-OCT are as follows. First, while measurement light 18 is focused onto the surface 17 of the sample 1 as an elongated shape, reference light is returned to the beam splitter 7, which is reflected by the reference mirror 8. In such an arrangement, the signal light 10 and reference light 11 are combined and imaged on the CCD 16. Thereupon, a fringe (spatial interference pattern) is produced on the surface of the CCD 16. The intensity of the fringe pattern is measured. Interferogram is calculated by Fourier transforms of the intensity of the fringe pattern with a computer. Further, the focusing/imaging of the measurement light and so forth is performed by two cylindrical lenses 19 that condense only in the x′ axis direction and one cylindrical lens 20 that condenses only in the y′ axis direction.
In the FD-OCT, because translation of the reference mirror 8 is unnecessary, the measurement time can be short. As an example, measurement time of about 150 msec has been reported. However, this method has following problems.
(Problem 1) The resolution in the transverse direction is low (transverse resolution; on the order of 100 μm).
When a spectral density function is calculated, it is assumed that the reflective faces within the sample extend to a fixed depth and, therefore, an accurate spectral density function is not obtained in a sample in which the depth of the reflective face changes abruptly in the transverse direction (y′ axis direction). Therefore, the resolution in the direction (y′ axis direction) parallel to the surface of the sample is poor, only values on the order of 100 μm having been reported.
(Problem 2) Narrow measurable range in the depth direction (measurable range; 6.0 mm)
The measurable range Lm in the depth direction is determined by the effective coherence length of each frequency component detected by the CCD. Suppose that the spectral width of each frequency component is Δf, and c is light speed, the measurable range Lm is given by Equation (1) (refer to the Equations appearing in Non-Patent Document 2). However, what is referred to here as the measurable range is not the measurable range in the depth direction of the sample but instead represents the measurable range according to the optical path difference between the light irradiated into the sample and the reference light. Therefore, the measurable range that appears in Non-Patent Document 2 is two times the measurable range in the depth direction of the sample.
                    (                  Equation          ⁢                                          ⁢          1                )                                                                                            L            m                                    =                  c                      Δ            ⁢                                                  ⁢            f                                              (        1        )            
In the FD-OCT, Δf depends on the pixel width of the CCD in the frequency axis direction (x axis). When an SLD of coherent length 34 μm is the light source, and a CCD with a pixel number in the frequency axis direction of 640 and a pixel interval of 13.3 μm is used, the measurable range calculated by means of Equation (1) is 9.0 mm (See Non-Patent Document 2). However, in moving away from zero on the y axis, the optical path length difference (OPD) between the signal light 10 and reference light 11 following division by the beam splitter 7 increases. As a result, when the fringe cycle approaches the pixel width of the CCD, averaging of the fringe is produced. As a result, the S/N drops and the range over which a clear interferogram can be calculated is reduced to ±6.0 mm for the OPD value (6.0 mm in the depth direction).
(Problem 3)
In measurement of a biological body, the intensity of the light that can be irradiated onto the sample is restricted. Therefore, efficient detection of signal light is important. However, in the FD-OCT, the signal light enters the photodetector (CCD) after a diffraction grating 21 and, therefore, there is the problem that a portion of the signal light is lost by the diffraction grating 21 and the detection efficiency of the signal light is poor.
(Problem 4)
Further, when detection using CCD is performed, the dynamic range representing the number of digits of the measurable intensity is no more than approximately 70 dB. Possibility of application of such detection to retina has been reported. However, that does not necessarily mean that such detection is sufficient for observation of a biological body.
(Problem 5)
Furthermore, there is also the problem that the measurement time is limited by the speed of the CCD, and there is a limit on increasing the measurement speed.
<B> Wavelength-Tunable Light Generator for Dental OCT and Dental OCT Device
The present invention described hereinbelow relates to wavelength-tunable light generator for dental OCT and dental OCT devices. It is extremely effective when applied to a cavity detecting device for diagnosis of the characteristics of a tooth by obtaining a tomographic image of a tooth. OCT is noninvasive to a biological body and has a high resolution. Therefore, applications not only to the tomography of retina but also to that of other organs have been tried (See Non-Patent Document 1, for example), and detecting the characteristics of a tooth, for example, may be considered (See Non-Patent Document 5, for example).
<C> Device for Measuring a Tomogram of Various Structures such as a Biological Body or Coated Surfaces
OCT is optical coherence tomography that is useful for imaging a tomogram of a retina or the like (See Non-Patent Document 7, for example). The OCT has attracted attention for being noninvasive to a biological body and for its high resolution. Applications to organs other than eye have also been tried (See Non-Patent Document 7, for example). One of the characteristics of this measurement method is its spatial resolution in the depth direction, and measurement devices with a resolution on the order of approximately 10 μm have been put to practical use. The resolution is determined by the spectral width of the light source. However, for OCT in practical use, to realize user-friendly, reliable, small-scaled and light-weighted systems, usually a semiconductor light-emitting element and, more specifically, a near-infrared SLD is used for the light source. That is, the resolution of OCT devices in practical use is limited by the spectral width of the SLD. The spatial resolution of OCT is inversely proportional to the spectral width of the light source and, therefore, the spectral width of the light source may be increased in order to increase the resolution. However, the spectral width of SLD is determined by the physical nature of the light-emitting layer and so forth and it is therefore difficult to increase the spectral width above the value realized now.
To overcome this limit, a trial, in which a plurality of SLDs of different center wavelengths are combined to substantially implement a broadband light source, has been proposed by Satou et al (See Non-Patent Document 8, for example).
[Patent Document 1] Japanese Patent Application Laid Open No. H6-53616
[Patent Document 2] Japanese Patent Application Laid Open No. H6-61578
[Patent Document 3] U.S. Pat. No. 4,896,325
[Patent Document 4] U.S. Pat. No. 3,471,788
[Patent Document 5] Patent Application No. 2003-335207
[Non-Patent Document 1] Chan Kin Pui, ‘Microscopic diagnostics using optical coherence tomography for clinical applications’, Optronics, Optronics Corp, Jul. 10, 2002, 247th Edition, pages 179 to 183
[Non-Patent Document 2] TERAMURA Yuichi, MIKUNI Masayuki, KAMINARI Fumihiko Proceeding of 23rd Meeting on Lightwave Sensing Technology, page 39
[Non-Patent Document 3] Handbook of Optical Coherence Tomography (edited by Brett E. Bouma, Guillermo J. Tearney), pages 364 to 367
[Non-Patent Document 4] YOSHIKUNI Yuzo ‘Developmental trends of wavelength-tunable lasers and expectations for system applications’, Applied Physics, Applied Physics Scientific society, 2002, 71st Volume, Eleventh edition, pages 1362 to 1366
[Non-Patent Document 5] Edited by Brett E. Bouma et al., Handbook of Optical Coherence Tomography, (USA), Marcel Dekker Inc., 2002, p. 591 to 612
[Non-Patent Document 6] CHOI, Donghak ‘High-speed, high-resolution OFDR-OCT using SSG-DBR laser’, Twenty-eighth Optical symposium lecture proceedings, Corp, Applied Physics Subcommittee meeting, Optical Society of Japan, Jun. 19, 2003, pages 39 to 40
[Non-Patent Document 7] Chan Kin Pui, ‘Microscopic diagnostics using optical coherence tomography for clinical applications’, Optronics, Optronics Corp, Jul. 10, 2002, 247th Edition, pages 179 to 183
[Non-Patent Document 8] Applied Optics February 2003, pages 7 to 11, SATOU, Manabu
[Non-Patent Document 9] Twenty-eighth Optical Symposium proceedings, Pages 39 to 40 (Published Jun. 19, 2003)